The present disclosure relates to ultrasonic imaging. More particularly, systems and methods for improved axial resolution and increased sensitivity associated with a harmonic ultrasound imaging modality are disclosed.
Ultrasonic imaging is used in many clinical applications because of its high image quality, safety, and low cost. Ultrasonic images are typically formed through the use of phased or linear-array transducers which are capable of transmitting and receiving pressure waves directed into a medium such as the human body. These ultrasonic transducers may be further assembled into a housing, which may contain control electronics, the combination of which forms an ultrasonic probe.
Ultrasonic probes are used along with transceivers to transmit and receive pressure waves through the various tissues of the body. The various ultrasonic responses are then processed by an ultrasonic-imaging system to display the various structures and tissues of the body.
Ultrasound imaging systems can create two-dimensional brightness or B-mode images of tissue in which the brightness of a pixel is based on the intensity of the received ultrasonic echoes. In another common imaging modality, typically known as color-flow imaging, the flow of blood or movement of tissue is observed. Color-flow imaging takes advantage of the Doppler effect to color-encode image displays. In color-flow imaging, the frequency shift of backscattered-ultrasound waves is used to measure the velocity of the backscatterers from tissues or blood. The frequency of sound waves reflecting from the inside of blood vessels, heart cavities, etc. is shifted in proportion to the velocity of the blood cells. The frequency of ultrasonic waves reflected from cells moving towards the transducer is positively shifted. Conversely, the frequency of ultrasonic reflections from cells moving away from the transducer is negatively shifted. The Doppler shift may be displayed using different colors to represent speed and direction of flow. To assist diagnosticians and operators, the color-flow image may be superimposed on the B-mode image.
Ultrasonic imaging can be particularly effective when used in conjunction with contrast agents. In contrast-agent imaging, gas or fluid filled micro-sphere contrast agents known as microbubbles are typically injected into a medium, normally the bloodstream. Due to their physical characteristics, contrast agents stand out in ultrasound examinations and therefore can be used as markers that identify the amount of blood flowing to or through the observed tissue. In particular, the contrast agents resonate in the presence of ultrasonic fields producing radial oscillations that can be easily detected and imaged. Normally, this response is imaged at the second harmonic, 2ft of the fundamental or transmit frequency, ft. By observing anatomical structures after introducing contrast agents, medical personnel can significantly enhance imaging capability for diagnosing the health of blood-filled tissues and blood-flow dynamics within a patient""s circulatory system. For example, contrast-agent imaging is especially effective in detecting myocardial boundaries, assessing micro-vascular blood flow, and detecting myocardial perfusion.
U.S. Pat. No. 5,410,516 to Uhlendorf et al. discloses that a radio-frequency (RF) filter can be used to selectively observe any integer harmonic (2nd, 3rd, etc.), subharmonic (e.g., xc2xd harmonic) or ultraharmonic (e.g., 3/2 harmonic) of ft to improve the microbubble to tissue signal ratio. The second harmonic has proven most useful due to the large bubble response at this frequency as compared to higher-order integer harmonics, subharmonics or ultraharmonics. The second harmonic also is most practical due to bandwidth limitations on the transducer (i.e.,  less than 70% bandwidth, where percent bandwidth is defined as the difference of the high-corner frequencyxe2x80x946 dB point from the low-corner frequencyxe2x80x946 dB point, divided by the center frequency.) However, single-pulse excitation techniques together with harmonic imaging suffer from poor microbubble-to-tissue signal-intensity ratios as large fundamental signals (ft) scattered from tissue mask the signals generated by the contrast agent.
As a result, of the discrimination problem associated with single-pulse excitation techniques, various multiple-pulse methodologies have been developed to suppress ultrasonic responses from anatomical tissues. These multiple-pulse excitation techniques result in diagnostic displays having an intensity that is responsive to the concentration of the contrast agent within the local insonified region.
Several techniques have been developed which take advantage of the primarily linear-response behavior of tissue to cancel or attenuate the linear-tissue signals. In several of these techniques, multiple transmit lines are fired along the same line of sight into the body. The transmit waveform is modified (e.g., in terms of power, phase, or polarity) from line to line to produce a variation in the response received by the transducer. These data points are then processed to remove the influence of their linear components to yield data that primarily contains the non-linear response of the contrast agents.
Although the above-described techniques work well in removing the influence of tissue generated signals, further improvements in resolution and system sensitivity are desired. Lateral resolution of a pulsed-echo ultrasound-imaging system depends on the ultrasonic-beam width. Axial resolution depends on the ultrasonic-pulse duration.
The lateral resolution may be improved by means of static or dynamic focusing using acoustic lenses or electronically focused transducer arrays. The width of the focused-ultrasonic beam is proportional to its wavelength or the frequency content of the transmit pulse.
The axial resolution may be improved by using high-frequency ultrasound or making the ultrasonic pulses shorter. However, high-frequency ultrasound pulses are limited in the depth of penetration due to tissue attenuation. Tissue attenuation increases with the frequency of the transmit pulses.
Generally, an ultrasound transducer is excited by an electronic waveform having a sharp voltage spike. In this case, the length of the transducer impulse response limits the duration of the ultrasonic pulse. Mechanical damping of the transducer further reduces the length of the impulse-response function. However, mechanical dampening sacrifices transducer bandwidth and sensitivity. Since it is common to use the same transducer for both transmitting ultrasonic pulses and receiving tissue-generated echoes, mechanical dampening is often an unacceptable solution as mechanical dampening generally limits transducer bandwidth.
To overcome the problems associated with high-frequency transmissions and the desire to shorten the length of the impulse response of the transducer, some have modified the excitation or transmit waveform that is applied to the transducer. The shape of the excitation waveform ultimately determines the shape and duration of the associated transmitted ultrasonic pulse.
Attempts to control the transmitted ultrasonic pulse waveform can be traced in the following patents. In U.S. Pat. No. 4,222,274 to Steven A. Johnson (1978), an apparatus is proposed that is capable of transmitting ultrasonic beams of two predetermined shapes. In U.S. Pat. No. 4,520,670 to Goran Salomonsson et al. (1982), a method and an apparatus is proposed for generating short-ultrasonic pulses by means of an excitation signal shaped as a weighted least-squares filter. Another system including a complex beamformer is described in U.S. Pat. No. 5,675,554 to Christopher R. Cole et al. (1996). The complex beamformer of Cole et al. is capable of producing focused-ultrasonic beams having a specified-carrier frequency and envelope.
Nevertheless, it is still desirable to be able to produce ultrasonic pulses having a variety of precisely-specified waveforms including those which can not be specified in terms of carrier frequency and envelope. Because of the limited available transducer bandwidth, the following problems arise for harmonic imaging. Because of the spectral falloff, both the transmitted and received signals have reduced sensitivity over the maximum-available transducer bandwidth. The limited transducer bandwidth also constrains both the transmitted and received-pulse frequency bandwidths to narrow ranges that fall within the available-transducer bandwidth when conventional-uncompensated transmission methods are used.
Furthermore, the transmitted pulses, when generated by a conventional means such as a tone-burst excitation, result in spectral sidelobes, which spill into the overlap region of the transmitted and received spectra. The sidelobes undesirably distort the response and the resultant image.
Moreover, the spectral-transducer amplitude and phase response distorts both the excitation transmissions and the received echoes so that the frequency spectrum associated with both the transmissions and the echoes are asymmetric.
A transmit-signal modifier reduces the aforementioned problems associated with limited-transducer bandwidth by compensating for the transducer response on transmit and/or receive and provides a method for realizing preferred-signal shapes for enhanced-harmonic imaging. A transmit-signal modifier may include a transmit controller having a digital-signal processor configured to calculate a drive spectrum that takes into account the impulse response of the ultrasound-transmit system. The digital-signal processor is further configured to determine a temporal-drive signal that results in the preferred transmit-spectrum shape when the drive signal is applied to the ultrasound transducer.
The digital-signal processor may employ inverse-Fourier transform methods including an inverse Fast-Fourier transform. The digital-signal processor may also employ alternative deconvolution methods to derive the transducer-compensated drive spectrum.
An echo-signal shaper may be realized with a receive filter having a center frequency at a designated harmonic of the transmit pulse center frequency. The receive filter may be adapted to a preferred receive-signal spectra and the impulse response of the ultrasound system including the transducer over the desired receive bandwidth. The required receive filter may be implemented by various deconvolution methods.
A method for enhancing the axial resolution and sensitivity of an ultrasound-imaging system is also disclosed. In its broadest terms, the method can be implemented by performing the following steps: selecting a preferred spectral shape of the acoustic ultrasound-transmit spectrum; identifying the impulse response of the transducer over the transmit spectrum; deriving a transmit-drive spectrum shape; determining the temporal-drive signal from the derived transmit-drive spectrum; and applying the temporal-drive signal.
The designated waveform will determine the frequency spectrum and the time duration of the transmitted-ultrasonic pulse thus enabling the use of the same transducer for different applications such as near field and far-field examinations. The precise shape of the transmitted ultrasonic-excitation signal can be selected to facilitate image-reconstruction techniques such as deconvolution or wavelet transform resulting in improved-axial resolution and superior-image quality. Similarly, by compensating for the adverse affects of the transducer-impulse response on the received echoes, a more accurate rendition of the tissues under observation can be attained.
Other features and advantages of the system and method for improved harmonic imaging will become apparent to one skilled in the art upon examination of the following drawings and detailed description. It is intended that all such additional features and advantages be included herein as protected by the accompanying claims.